System for quantitative radiographic imaging

ABSTRACT

A system for spectroscopic imaging of bodily tissue in which a scintillation screen and a charged coupled device (CCD) are used to accurately image selected tissue. An x-ray source generates x-rays which pass through a region of a subject&#39;s body, forming an x-ray image which reaches the scintillation screen. The scintillation screen reradiates a spatial intensity pattern corresponding to the image, the pattern being detected by a CCD sensor. The image is digitized by the sensor and processed by a controller before being stored as an electronic image. Each image is directed onto an associated respective CCD or amorphous silicon detector to generate individual electronic representations of the separate images.

RELATED APPLICATIONS

This application is a continuation-in-part of U.S. patent applicationSer. No. 08/438,800 filed May 11, 1995 which is a Continuation-in-Partof U.S. patent application Ser. No. 07/853,775 filed Jun. 2, 1992 nowU.S. Pat. No. 5,465,284, which is the U.S. National Phase ofInternational Application No. PCT/US90/07178, filed Dec. 5, 1990 andwhich is a continuation-in-part of U.S. patent application Ser. No.07/446,472, filed Dec. 5, 1989, now U.S. Pat. No. 5,150,394, all of theabove applications being incorporated herein by reference in theirentirety.

BACKGROUND OF THE INVENTION

In recent years the use of radiological examining equipment to makemeasurements of bone density in patients has continually increased. Inparticular, the use of such equipment in diagnosing and analyzingosteoporosis has become prevalent in the medical community. Osteoporosisis characterized by the gradual loss of bone mineral content or atrophyof skeletal tissue, resulting in a corresponding overall decrease inaverage bone density. Such a condition is common in elderly women andgreatly increases the risk of fracture or similar bone related injury.

The presently available techniques for the radiological measurement ofbone density utilize a rectilinear scanning approach. In such anapproach, a radiation source, such as a radionuclide source or an x-raytube, and a point detector are scanned over a patient in a rasterfashion. This scan results in an image which has been derived from thepoint-by-point transmission of the radiation beam through the bone andsoft tissue of a patient. The calculation of the bone-mineralconcentration patient. The calculation of the bone-mineral concentration(the "bone density") is usually performed by a dual energy approach.

The current rectilinear scanning approach is generally limited by itslong scanning time and its lack of good spatial resolution. The poorspatial resolution results in an inability to provide an imagedisplaying high anatomical detail and which will permit accuratedetermination of the area in the scan occupied by bone. Moreover, theoutput of the x-ray source and the response of the detector must beclosely monitored in order to assure high accuracy and precision.

SUMMARY OF THE INVENTION

In accordance with the present invention, a stationary bone densitometryapparatus is provided for examining a subject's body. A dual energyx-ray source directs a beam of x-ray radiation toward the subject'sbody. The radiation is applied to the entire region of the body beingexamined. A scintillation screen receives the x-ray radiation passingthrough the body of the subject, and emits radiation in the visiblespectrum with a spatial intensity pattern proportional to the spatialintensity pattern of the received x-ray radiation.

A charge coupled device (CCD) then receives radiation from thescintillation screen. This CCD sensor generates a discrete electronicrepresentation of the spatial intensity pattern of the radiation emittedfrom the scintillation screen. A focusing element between the screen andthe CCD sensor focuses the scintillation screen radiation onto the CCDsensor. To prevent ambient radiation from reaching the CCD sensor, thepresent embodiment employs a shade or hood surrounding a region betweenthe scintillation screen and the CCD sensor. A CCD controller thenprocesses the electronic representation generated by the CCD sensor, andoutputs corresponding image data.

A dual photon x-ray source is used to allow the examination to beperformed with x-rays at two different energy levels. This source can bean x-ray tube, or a radionuclide source with a filter element to removeone of the energy levels when desired. Correlation of the image dataretrieved using each of the two x-ray energy levels providesquantitative bone density information.

A focusing element between the scintillation screen and the CCD sensorcan take the form of a lens or a fiber optic reducer. An imageintensifier can be used in conjunction with the CCD sensor. The imageintensifier can be a "proximity type" image diode or a microchannelbased device. It can also be directly attached to the CCD. An imagestore used with the CCD controller allows manipulation of the CCD sensoroutput signals by a data processor. This includes the correlation ofmeasurements utilizing x-ray beams of two different energy levels. Thesystem can also be adapted to operate at higher shutter speeds enablingthe counting of x-ray transmissions. This provides energy measurementsof x-ray transmissions that are useful in certain applications.

In an alternative embodiment, a detector made of amorphous silicon isused to receive and detect the radiation from the scintillation screento generate the electronic representation of the spatial intensitypattern of the x-ray pattern. The amorphous silicon detector can replacethe CCD detector or it can be used to receive the x-rays directly.

In another preferred embodiment, the apparatus of the invention includestwo scintillation screens, each of which is associated with its ownrespective CCD detector or amorphous silicon detector. One of thescintillators is reactive to high-energy x-rays and generates an opticalimage of the spatial intensity pattern of the high-energy x-ray pattern.Its associated detector detects the image and generates an electronicrepresentation of the high-energy x-ray pattern. The other scintillatoris reactive to low-energy x-rays to simultaneously generate an opticalimage of the low-energy pattern. Its associated detector generates anelectronic representation of the low-energy x-ray pattern. The dataprocessor performs the correlations of the measurements for the x-raysat two different energy levels.

An additional preferred embodiment is directed to systems and methods ofimaging spectroscopy where charge coupled device (CCD) is opticallycoupled to a scintillator and measures or counts the spatial intensitydistribution of a radionuclide that has been introduced into bodilytissue, either in vivo or in vitro. CCD's of sufficient thickness can beused to measure gamma ray events without the use of a scintillator incertain applications. The CCD has sufficient resolution and sensitivityto measure such distributions accurately, usually in less than twominutes. Radiation sources that emit radiation having an energy in arange between 10 and 2,000 keV, and preferably in the range between 20and 600 keV, are delivered to the cancerous tissue or any other suitablepathologic abnormality.

The CCD acquires "frames" of information by counting the number ofgamma-ray events over a selected period of time. Each frame, or asequence of frames that have been added or summed to provide an image,can be filtered using pulse height analysis techniques to substantiallyreduce or eliminate scattered radiation. Pulse height analysis can alsobe utilized to discriminate between signals having different energylevels that contain diagnostically significant information. The system'sdiscrimination and energy measuring capabilities render it suitable fordiverse applications.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a perspective view of the imaging system the presentinvention.

FIG. 2 illustrates in schematic view a bone density measuring apparatususing a lens to focus image data from a scintillation onto a CCD sensor.

FIG. 3 illustrates in schematic view a bone density measuring apparatususing a fiber optic reducer to deliver an image from a scintillationscreen to a CCD sensor.

FIG. 4 illustrates another preferred embodiment for the scintillationscreen employing a fiber optic plate.

FIG. 5 is an illustration of the pixel array of a binnable CCD sensor.

FIG. 6 is an alternative preferred embodiment to the bone densitymeasuring apparatus of FIG. 2.

FIG. 7 is another alternative preferred embodiment to the bone densitymeasuring apparatus of FIG. 2.

FIG. 8 is a perspective view of a scanning system of the presentinvention.

FIG. 9 is a schematic sectional view illustrating the sensor controlsystem.

FIG. 10 is a schematic sectional view illustrating a frame transfer CCDused for both emission and transmission studies.

FIG. 11 is a schematic sectional view of a CCD imaging system for bothemission and transmission studies.

FIG. 12 illustrating a process flow sequence that is used in performingthe imaging methods of the present invention.

FIG. 13 is an alternate embodiment of a CCD imaging system that can beemployed for both emission and transmission studies.

FIGS. 14A and 14B illustrate a process flow sequence for conductingemission and transmission studies of tissue.

FIG. 15 is a schematic diagram of an alternative preferred embodiment tothe bone densitometry measuring apparatus of FIG. 2 using dualscintillation screens and dual detectors.

FIG. 16 is a schematic diagram of an alternative preferred embodiment tothe bone densitometry measuring apparatus of FIG. 15.

FIG. 17 is a schematic diagram of a variation of the bone densitometrymeasuring apparatus of FIG. 16.

FIG. 18 is a schematic diagram of another alternative preferredembodiment to the bone densitometry measuring apparatus of FIG. 15having dual amorphous silicon image sensors.

FIG. 19 is a schematic diagram of an alternative detection structureincluding dual amorphous silicon image sensors which can be used withthe various embodiments of the bone densitometry measuring apparatus ofthe invention.

FIG. 20 is another preferred embodiment of a bone densitometer forstatic, scanning or stepped imaging procedures.

FIGS. 21A-21C illustrate alternate embodiments for the detector assemblyof FIG. 20.

FIGS. 22A-22B illustrate scanning or stepped imaging procedures.

FIG. 23 is another preferred embodiment of an imaging system inaccordance with the invention.

FIG. 24 is another preferred embodiment of an x-ray imaging system inaccordance with the invention.

FIG. 25 is a schematic diagram of a detection structure used for bonedensitometry measurements and tissue lesion imaging in accordance withthe present invention.

FIG. 26 is a preferred embodiment of the invention in which a dualspaced array is used for digital mammographic imaging and quantitativeanalysis.

FIG. 27 illustrates another preferred embodiment in which the imagingelements in each linear array positioned at a different angle relativeto the patient and the x-ray source.

FIG. 28 illustrates a preferred embodiment in which a large number ofimaging elements are arranged in a large array conforming in size to astandard x-ray film cassette which can have three or more spaced lineararrays.

FIG. 29 illustrates a cross-sectional view of a linear array using acommon fiber-optic plate and scintillator.

FIG. 30 illustrates a system for translating the array relative to aradiation source.

FIGS. 31A and 31B illustrate the process of a two step imaging sequence.

FIG. 32 illustrates a system for sequential imaging or scanning oftissue in which the array is moved relative to a source.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

In FIG. 1 a preferred embodiment of the invention for performing bonedensitometry studies uses a detector 10 and either an x-ray tube 12 or aradionuclide radiation source such as Gadolinium-135. The detector 10comprises a scintillating plate 20 which is optically coupled to atwo-dimensional charge-coupled device 24 (CCD). The CCD is a twodimensional array of detectors integrated into a single compactelectronic chip. The optical coupling between the scintillating plate 20and the CCD 24 is accomplished by an optical grade lens 25. Such a lensshould have a low f-number (0.6-1.8) for adequate light collection fromthe screen. The collection efficiency (E) of light from thescintillating plate emitted in the direction of the CCD can be computedby the equation: ##EQU1## where: t: Transmission factor of light throughthe lens

m: magnification from the Scintillating plate to the CCD

f: f-number of the lens

In an alternate approach, the optical coupling between the scintillatingplate and the CCD can be performed with a fiber optic reducer.

Referring to FIG. 2, a bone densitometry apparatus 10 has an x-ray tube12 which delivers a beam of x-rays 14 towards the body of a subject 16being examined. The x-ray tube is capable of emitting x-ray radiation ateach of two distinct energy levels. The two energy levels are used toobtain two distinct x-ray images of the patient, as is discussed later.Note in comparison to FIG. 1, the source can be positioned above thepatient and the detector below the table.

When the subject 16 is irradiated with the x-ray energy, a percentage ofthe x-rays reaching the subject 16 is absorbed by the subject's body,the amount of absorption depending on the density of bone or tissue uponwhich the x-rays are incident. Since x-rays generally travel in astraight line, the x-ray energy exiting the subject's body on the sideof the body away from the source 12 is a spatial representation ofabsorption in the subject's body, and therefore of relative tissue andskeletal densities.

To receive the x-rays passing through the subject's body, ascintillation screen 20 is provided on the side of the patient away fromthe x-ray source 12. The scintillation screen 20 is a fluorescentmaterial sensitive to x-rays, and when it receives x-ray energy itreradiates visible light. The spatial intensity patterns of theradiation emitted from the scintillation screen is proportional to thespatial intensity pattern of the x-ray radiation received by the screen20. Thus the scintillation screen 20 provides an image in the visiblespectrum, or alternating in the ultraviolet or near infrared, which isregionally proportional to the x-ray image reaching the scintillationscreen 20.

A lens 22 is positioned between the scintillation screen 20 and a CCDsensor 24. The CCD sensor 24 is an array of photosensitive pixels usingclosely spaced MOS diodes which convert photons to electrons and therebygenerate a discrete electronic representation of a received opticalimage. The lens 22 faces the scintillation screen and focuses thevisible light emitted from the scintillation screen 20 through the lens22 and onto the surface of the CCD sensor 24. In order to preventambient light from reaching the CCD sensor, a shade surrounding theregion between the scintillation screen 20 and the lens 22 is providedin the form of a photographic bellows 26. The shading of bellows 26serves to reduce the optical noise level of the image signal reachingthe CCD sensor 24.

Although the scintillation screen 20 absorbs most of the x-rays incidentupon it, some may still be transmitted through the screen 20 andinterfere with the optical image signal of the scintillation screen 20.The direct interaction of x-rays with a CCD sensor produces very brightpixels resulting in a "snow" effect in an optical image detected by thesensor. In addition, prolonged direct x-ray irradiation of a CCD sensorcan increase its dark current. For these reasons, an optical gradelead-glass or lead acrylic filter 28 is positioned between thescintillation screen 20 and the lens 22 or alternatively, between thelens and the CCD. The lead-glass filter 20 absorbs most of the strayx-rays and prevents them from reaching the CCD sensor 24. Ananti-scatter grid 29 is used between the patient and scintillationscreen for preventing scattered x-rays from reaching the screen.

During a typical examination, the subject 16 is placed between the x-raysource 12 and the scintillation screen 20. The x-ray source is thenactivated for a short time interval, typically one to five seconds. Asx-rays are differentially transmitted and absorbed through the body ofthe subject 16, they interact with the scintillation screen 20. Uponinteraction, the screen 20 emits light in the visible part of theelectromagnetic spectrum. In the present embodiment, the scintillationscreen is a terbium-activated material and emits light in the region of540 nm.

The light emitted from the scintillator is transported to the CCD sensorvia the lens 22. Upon interaction with the CCD sensor 24, light energyis converted into electrons which are stored in each pixel of the CCDsensor 24. The CCD sensor 24 of the present embodiment consists of512×512 pixels, but such sensors come in a number of different sizes.The CCD sensor "integrates" the image signal from the scintillationscreen in that it senses the optical image and stores charge during theentire x-ray exposure interval. After termination of the x-ray exposure,the discrete representation in the CCD 24 is read out by CCD controller30. The CCD controller 30 reads the image representation from the CCDsensor 24 pixel by pixel and organizes it into a digital array. Thedigital array, representing spatial position and x-ray intensity, isthen output to a memory or image store 32. From the image store 32, theimage can be accessed by a data processor 34 for performing imageprocessing techniques. A cathode ray tube (CRT) 36 is also provided toallow the image to be displayed before or after processing by dataprocessor 34.

Unlike other conventional detection schemes, such as film screenradiography, CCD-based imaging provides a linear quantitativerelationship between the transmitted x-ray intensity and the chargegenerated in each pixel of the CCD. After the first high energy x-rayexposure is acquired, the resulting image is stored in image store 32and a second exposure with a low energy x-ray beam is acquired with thesubject 16 in the same position. During this exposure, a low energyx-ray beam is used which is typically at about 70 kVp with a tube.current at about 1 mA. The tube is capable of accelerating electrons at40 kVp and up to approximately 140 kVp. Note that the tube potential andcurrent are controlled by the computer menu. The low energy x-ray imageis then stored in image store 32 with the high energy exposure. Eachimage provides quantitative information about the relative transmissionof x-rays through soft tissue and bone.

Once both images are obtained, comparative processing techniques of dualphoton absorptiometry are applied to determine quantitative densitymeasurements of those body regions scanned by the x-rays. Thecorrelation of two images generated by x-rays of two different energylevels over a short time interval results in the substantial reductionin the likelihood of systematic pixel-by-pixel errors caused byinstability of the x-ray tube output.

Because the present embodiment of the invention is concerned with anarea detector as opposed to a scanning detector, the measurement timenecessary for a densitometry examination is greatly reduced. Rather thanscanning across the region to be examined in a rectilinear fashion, theentire region is irradiated simultaneously and the resulting imageprocessed simultaneously. Typically, the entire procedure using thepresent dual photon technique lasts 30 to 60 seconds, depending on thepower of the x-ray tube and processing speed of the supportingelectronics.

FIG. 3 shows an alternative embodiment to that of FIG. 2. In thisembodiment, the x-ray tube source 12 of FIG. 2 is replaced withradionuclide source 40. The radionuclide source is gadolinium-135.Gadolinium-153 emits photons simultaneously in two energy bands, a lowerenergy band of 44 keV and an upper energy band of 100 keV. Thus, thegadolinium source is a dual photon radiation source. In order to allowthe images from the two different energy levels to be obtainedseparately, an x-ray filter 42 is placed between the source 40 and thesubject 16. In the present embodiment, the filter 42 is copper or aK-edge filter, and eliminates nearly all of the low energy (44 keV)emission from the beam. Removal of the filter restores the beam to itsdual energy nature. The filter 42 is implemented as an electromagneticshutter which may be opened and closed in the line of the x-ray beam. Ahigh energy image is acquired first with the filter shutter closed,after which an image is obtained using the dual energy beam with theshutter open.

Both electronic images are stored, and an image representative of thetransmission of only the low energy photons is obtained byelectronically subtracting the high energy image from the dual energyimage with the data processor 34. Once both images are obtained,comparative dual photon processing techniques are used to makequantitative density calculations.

An additional feature of the embodiment of FIG. 3 is the replacement ofthe lens 22 of the FIG. 2 embodiment with a fiber optic reducer 44. Thefiber optic reducer 44 is a focusing device consisting of a large arrayof optical fibers packed tightly together, and leading from thescintillating screen 20 to the CCD sensor 24. Near the CCD sensor 24,many of the fibers can be fused together, thus combining the signalspresent on individual fibers. The effect is a compression of the imagefrom the input of the reducer 44 at the scintillation screen 20 to thereducer output at the CCD sensor 24. In this manner, the reducer 44effectively focuses light from the scintillating screen 20 onto the CCDsensor 24 without the necessity of a lens for the focusing region.

Although they are shown together in FIG. 3, it is not necessary to usethe fiber optic reducer 44 with the radionuclide source 40. Eitherelement can be substituted into the configuration of FIG. 2individually. The x-ray filter 42, however, should be used with theradionuclide source 40 to provide a dual photon discriminationcapability. Note, however, that pulse height analysis can be performedin conjunction with the embodiment of FIGS. 10 & 11.

FIG. 4 shows an alternative to the scintillation screen 20 of FIGS. 2and 3. The screen 48 depicted by FIG. 3 is a scintillating fiber opticplate. The plate 48 is a fiber optic faceplate consisting ofscintillating fibers 50 running through the plate. The fiber optic plateis optically interfaced to the CCD in essentially the same way as thescintillation screen 20 of FIG. 2, but the fiber optic plate 48 allowsfor greater quantum efficiency due to increased x-ray stoppingcapability.

Shown in FIG. 5 is a representation of the pixel array of the CCD sensor24. The array shown in FIG. 5 is only 10×10 for illustrative purposes,and the actual array can be of different dimension. Each pixel in thearray is an individual photosensitive element which contributes to theoverall image detected by the array. A feature of the CCD sensor of thepresent embodiment is the capability of the pixels of the sensor 24 tobe "binned" together. The binning of the pixel array refers to theability of the sensor electronics to combine groups of pixels togetherto form "super pixels" which are then identified as single pictureelements.

Charge is binned by combining charge packets contained in two or moreadjacent potential wells into a single potential well during chargereadout. Serial and parallel binning can be combined to perform twodimensional binning from any rectangular group of wells or detectorelements.

The dark lines in the binnable array of FIG. 5 illustrate whereindividual pixels might be grouped together. For example, the four upperleft hand corner pixels 50 can be binned together through control of theCCD sensor 24 to form a super pixel. The super pixel is then identifiedby the CCD electronics as a single pixel, the light intensity reachingeach pixel 50 being averaged across the surface of the entire superpixel. In this manner, the dimension of the array can be electronicallycontrolled. As can be seen in FIG. 5, if groups of four pixels arebinned together across the 10×10 array, the overall array dimensionbecomes 5×5. Although the binning of the CCD sensor 24 reduces theresolution of the pixel array, the relative percentage of noise is alsoreduced, thus providing an improved signal to noise ratio.

The following x-ray data acquisition approach is an alternative to theone described above. In this approach, an image is acquired at highenergy and the CCD is read in the normal non-binned mode. Due to thehigh penetration of the high energy beam through the body, the x-rayfluence exiting the body is high as compared to that of the low energybeam. Therefore, the resulting charge signal per CCD pixel is relativelystrong. This image is stored as the high energy image. Also, this imageis used in order to compute the area of the bone to be measured bymanual selection of the region of interest or by automatic edgedetection. Therefore, we take advantage of the high resolution image forgreater accuracy in the measured bone area. Previously, the accuracy andprecision of bone density measurements are limited to a great extent bysuboptimal spatial resolution. The next image which is acquired with lowenergy is read out by the pixel binning approach, e.g., using a 2×2pixel binning. The transmission of the low energy beam through the bodyis low as compared to the high energy beam. Therefore, in order torecord a strong signal in each CCD pixel we must increase the radiationdose.

Alternatively, the binning technique can be used for the low energy inorder to increase the signal to noise ratio and to a decrease theradiation dose. This dual mode acquisition procedure is a very powerfultool for improving the signal to noise ratio and lowering the radiationdose to the patient.

Although the arrangement of optical elements as shown in FIGS. 2 and 3represent preferred embodiments, the functionality of the system is notdependent upon such an in-line type of optical transmission. FIG. 6shows an alternative arrangement of optical elements where the CCDsensor 24 is set at an angle relative to scintillation screen 20, andmirror 52 is used to reflect the radiation given off by thescintillation screen toward the CCD sensor 24. Lens 22 is shown betweenCCD sensor 24 and mirror 52 and focuses the image onto the CCD sensor.However, the focusing of the scintillation screen image can take placebefore or after the image reaches mirror 52. In fact, the mirror itselfmay be shaped to provide focusing of the image from the scintillationscreen 20.

FIG. 7 shows another alternative arrangement of optical components. InFIG. 7 the subject 16 is suspended by a support 54 which is transparentto x-rays. The support 54 keeps the subject 16 elevated a distance abovescintillation screen 20. As the x-rays reach scintillation screen 20,the screen 20 reradiates image data from the same surface upon which thex-ray radiation is incident. Mirror 52 is now aligned to reflect thisimage towards CCD sensor 24 which collects the image as focused throughlens 22 to be processed by the CCD controller 30.

As with the arrangement of FIG. 6, the focusing of the image from thescintillation screen 20 may take place before or after it is reflectedby the mirror 52, or may be focused by the mirror 52 itself. Inaddition, any of the optional elements previously discussed may besubstituted into the arrangement of FIG. 5 or FIG. 7. This includes thex-ray absorbing screen 28, the anti-scatter grid, the fiber opticreducer 44, and the fiber optic faceplate 48.

A very effective, radiation dose-efficient approach for reducing x-rayscatter and increasing the dynamic range of electronically acquiredx-ray images is the use of a slit-scan method. In this approach, a fanbeam of x-rays is scanned over the patient and a linear array ofdetectors is used to detect the transmitted radiation. In typicalapplications the length of the detector restricts the width of the areathat can be covered with one pass. Also, many small linear CCD orphotodiode arrays are used to form a line of detection. This results ina rather complex detector assembly. If cooling of the detector assemblyis required, it is difficult to accomplish for such an extendeddetector. Also, image intensification by using an electronic intensifierbecomes difficult and very costly.

An alternative embodiment for dual energy bone densitometry takesadvantage of the merits of slit-scan geometry without using a linear CCDor photodiode array. This approach is illustrated schematically in FIG.8. An area CCD sensor 64 is used in conjunction with a line-to-areafiber optic converter 62. This converter can be made of flexible orrigid optical fibers with cladding of lower index of refraction than thecore material. As shown in FIG. 8, the CCD 64 is divided into a numberof rows and a fiber optic ribbon is optically coupled or bonded to eachrow. The coupling of the CCD 64 to the converter 62 can be accomplishedusing the various systems described in connection with otherembodiments. An extramural absorber can be used to prevent lightcrossing from one fiber to another. The other ends of each ribbon arearranged in tandem to form a linear sensor. In front of the linearsensor (input end), an x-ray converting scintillator 60 is used such asgadolinium oxysulfide activated with terbium (GOS:Tb). Alternatively, ascintillating fiber optic plate can be used for improved quantumefficiency at higher energies. A linear x-ray sensor with a very compactarea detector is employed with the slit-scan embodiment.

A typical linear detector of this type comprises a few ribbons in tandemalong the length of the slit, and from one to a multitude of ribbonsacross the width of the detector slit.

In a typical example, consider a 512×512 pixel CCD where each pixel hasan area of 20×20 microns. A fiber optic bundle with individual fibers of60 microns in diameter is used for the embodiment. On the CCD each fiberwill cover an area of approximately 3×3 pixels. Perfect alignmentbetween each pixel and fiber is desirable but it is not essential forthis application. Close packing of the fibers will result in an array of170×170 or a total of 29,127 fibers covering the entire area of the CCD.Each ribbon of fibers corresponds to one row consisting of 170 fibersand covering approximately 512×3 pixels on the CCD. If all ribbonsemerging from the CCD were arranged in tandem, the linear sensor wouldbe approximately 175 cm in length. Alternatively, the ribbons can bearranged with a small number in tandem and a small number across thewidth of the slit. Using the above CCD, a 15.3 cm linear detector can bemade with approximately 15 ribbons in tandem thus using only a smallfraction of the CCD area.

Full use of the CCD area can be made by stacking the ribbons in groupsof 15, (one ribbon per CCD row), thus creating a quasilinear detectorconsisting of an array of 2,550×11 fibers optically coupled to an x-rayscintillator. The dimensions of this slit detector will be 153×0.66 mmwith a total sensing area of 1.0 cm². It is important to note that thetotal sensing area of the slit must be approximately equal to the totalarea of the CCD and the linear dimensions of the fiber optic output mustbe approximately the same as the linear dimensions of the CCD. A wideror longer slit will result in a larger area at the output end. In thiscase, a larger CCD can be used or a fiber optic reducer optically bondedbetween the fiber optic converter and the CCD. Alternatively, theconverter itself can be tapered to match the size of the CCD. For higherspatial resolution the fiber optic converter is made with optical fibersof smaller diameter (5-6 microns).

If higher signal amplification is required for some high detail low doseapplications, a proximity focused image intensifier can be opticallybonded between the fiber optic taper and CCD or between the fiber opticconverter and fiber optic taper. The image intensifier can be aproximity diode type or a microchannel plate device, both commerciallyavailable. Alternatively, an integral assembly of CCD and intensifiercan be used commonly called an "intensified CCD". Another approach is touse a lens coupling between the output surface of the fiber opticconverter and the intensified or non-intensified CCD.

Cooling of the CCD can be accomplished easily by a thermoelectriccooler. Cooling is required only when very high contrast resolution isrequired and the image acquisition time is relatively long. If the CCDis read out at 500 kHz (5×10⁵ pixels/sec), an area of 150 mm×150 mm ofthe subject can be scanned in approximately 114 seconds (approximately 2minutes). Faster scanning is attainable by increasing the readout rateof the CCD.

Alternatively, a frame transfer CCD such as the one illustrated in FIG.10 can be used for faster scanning. This device uses one half of itssensing area for storage and not for sensing. In this way the transferof the image from the sensing area 91 to the storage area 93 isaccomplished in a few milliseconds. A smaller CCD such as a 128×128 or a64×64 element could be used for this purpose in a similar arrangement aswith the 512×512 CCD. Also, larger area CCDs can be used for thispurpose. Pixel binning as described previously can be applied in thisdetection approach. A Gadolinium-135 (Gd-153) radiation source can beused as described in previous sections in place of an x-ray tube. TheGd-153 source is a small pellet or a collimated line source parallelwith the long dimension of the detector.

The line to area conversion design enables us to remove the CCD from thedirect path of the x-ray beam, thus it allows for easy shielding of theCCD from direct x-ray interactions. This prolongs the useful life of theCCD and it alleviates the "snow" effect which results from directinteractions of x-rays with the sensor. Moreover, this approach allowsfor greater light transport efficiency between the scintillator and CCDthan lenses or fiber optic tapers. Note that the pixel binning approachenables the operator to select the desired spatial resolution andcontrast without any mechanical modifications on either the x-ray beamor the detector collimator. The pixel size of the detector whichdetermines resolution and contrast can be controlled by a command fromthe computer. This x-ray imaging modality can be used very effectivelyto optimize the scan depending on patient size, and medical history.

An alternate approach provides an improved rectilinear scanning methodfor quantitative x-ray radiography. In this embodiment, a twodimensional CCD optically coupled to a scintillator is used as thedetector of x-rays in a rectilinear scanning mode. The CCD may be a fullframe or a frame transfer device. The frame transfer CCD will enablefaster data scanning and acquisition.

The CCD scintillator assembly is extremely critical to the performanceof the system. Direct optical bonding of a polycrystalline scintillatorsuch as gadolinium oxysulfide with the CCD is possible but this approachis not efficient in shielding the CCD from direct x-ray interactions. Ifthe thickness of the layer is increased the spatial resolution of thex-ray images degrades due to light diffusion. The use of a scintillatingfiber optic plate between the polycrystalline scintillator and the CCDprovides a solution to this problem.

A scintillating fiber optic plate is a fiber optic faceplate designed toconvert x-rays or U.V. light into green light with peak emission atabout 550 nm. This faceplate is manufactured with extra mural absorberto prevent light diffusion between individual fibers. The area of thescintillating fiber optic plate must cover the CCD completely. Thedesirable thickness depends on the energy of the x-ray radiation. Athickness of 5 to 10 mm is preferable but a thinner or thicker plate canbe used. The use of a very thick scintillating fiber optic plate such as10 mm or 20 mm will eliminate virtually any undesirable direct x-rayinteractions with the CCD. The scintillating fiber optic plate can alsobe used without the thin layer phosphor. However, the combination of thetwo will produce better image quality at a reduced radiation dose to thepatient. Alternatively a conventional fiber optic plate can be used as asubstrate to the scintillating fiber optic plate. The optical couplingof the polycrystalline phosphor on the fiber optic can be accomplishedby direct deposition techniques or by using an optical adhesive.

In an alternate approach, a bent fiber optic bundle can be used betweenthe scintillator and the CCD. The geometry of the bent bundle allows forextremely effective shielding of the CCD from extraneous x-rayradiation. A lens coupling between the CCD and the fiber optic convertercan also be used. For improved sensitivity, a proximity focused imageintensifier, an image diode or microchannel plate can be used at theinput end of the fiber optic or between the fiber optic bundle and theCCD. A preferred approach is to use the intensifier at the input end. Ascintillator can be optically bonded to the input of the intensifier oran intensifier with a scintillating fiber optic input plate can be used.

The x-ray tube is aligned in a C-arm configuration with the detector.The x-ray beam is approximately congruent with the area of the detectorwhich is approximately 1×1 cm at the detector plane. As x-rays aretransmitted through the patient, some (20%-60%) are absorbed by theprimary polycrystalline scintillator producing visible light. This lightis transmitted through the optically transparent fiber optic faceplatein the direction of the CCD. The x-rays not interacting with the primaryscintillator will be absorbed by the fiber optic faceplate. If ascintillating fiber optic faceplate is used, these x-rays will beabsorbed in the fibers thus producing additional scintillations.Therefore, the scintillating fiber optic plate acts as a lightconduction device, x-ray shield, secondary x-ray detector and an x-raysignal amplifier.

Upon interaction of the x-ray induced light with the photosensitivesurface of the CCD an electron charge is generated which is proportionalto the number of x-ray interactions in the scintillators. The cumulatedcharge on the CCD is then read out. However, in this rectilinearscanning mode, each CCD readout will correspond to a small segment ofthe total image, approximately one square centimeter. Therefore, theentire image is acquired by spatial additional of each image segment.For example, if a 15×15 cm field is covered and the sensor area is1.0×1.0 cm, 15² (225) segments must be acquired and synthesized. A512×512 pixel CCD operating at 500 kHz will read out each segment in 0.5seconds and will require about 2 minutes for the entire scan at a scanspeed of about 2 cm/sec. Faster scanning is attainable by increasingboth the scanning speed and the readout rate of the CCD.

A dual-energy scan will be acquired by first scanning the entire area athigh tube potential, typically 130 kVp without binning and thenrepeating the scan at low tube potential, typically at about 70 kVp withbinning. An automatic slide mechanism places high aluminum filtrationfor the high energy beam and less filtration for the low energy beam asdescribed previously. The images of each energy level are stored in thecomputer for subsequent dual photon analysis. Pixel binned acquisitionwill be possible at both energies for improved precision. Where bothhigh and low energy images are identically binned, this produces anexact correlation between the images produced. A third high energy-highresolution image can then be used to define the outline of the objectbeing scanned. Note that a gadolinium isotope source with a shutter canbe used.

Alternatively, the energy level of the tube can be switched from low tohigh for each segment of the acquisition and each segment representinghigh and low energy is stored for subsequent analysis.

An alternate approach employs light intensification from the screen tothe CCD sensor. In this approach, an electrostatically focused imageintensifier (In FIG. 2) is employed as the primary detector in place ofthe scintillating plate. This intensifier preferably employs Cesiumiodide input phosphor with an approximate diameter of 15 cm andthickness of 0.3-0.5 min. The high voltage of the image intensifier tubecan be reduced to approximately half the normal value. A reduction inthe image intensifier accelerating potential will contribute to animprovement in the image contrast characteristics and dynamic range ofthe device. The CCD sensor is optically coupled to the output phosphorof the image intensifier by a fast lens with an f-number of about 1:1.0.Due to the high signal intensification, cooling of the CCD is notessential but it can be applied if very low thermal noise levels aredesirable. The use of an intensifier allows for the use of a CCD withlower noise performance characteristics, thus lowering the cost andcomplexity of the instrument.

Ideally, the detected signal is produced by x-rays that have beentransmitted through the body without any scatter interaction. Detectionof large amount of scatter events will result in non-linearities and ina reduction in the dynamic range. Effective suppression of scatter isaccomplished by using a small field of view, typically 10 cm×10 cm andby using a air gap (approximately 20 cm) between the patient and thescintillating plate. Alternatively a small field of view can be used inconjunction with a linear or crossed antiscatter grid.

An internal instrument stability control system has been incorporated toprovide a means of automatic compensation for any instabilities in thex-ray tube potential and current. The stability control device is notessential for the operation of any of the described techniques but itprovides better reliability and precision in the measurement of bonedensity. A schematic representation of the proposed device is shown inFIG. 9. The output of the x-ray tube 12 is monitored by a pair of x-raysensors 70 placed at a secondary x-ray beam port 78 adjacent to the mainbeam port 80 near the tube window. The sensors can be silicon diodes,cadmium telluride radiation sensors or any other solid state x-raysensor. Alternatively, a pair of compact photomultiplier-scintillatorsor a photodiode scintillator assembly could be used. Both detectorsoperate in the charge integration mode and the detected signal iscontinuously monitored as a function of time during the entire' scan foreach energy. This time varying signal is digitized and stored in thecomputer memory. The change in the filtration of the secondary beam withenergy is identical to that in the main beam because it is controlled bythe same filter changing mechanism. As described further in connectionwith FIG. 12 the sensor system can be used to normalize the detectedinformation or to control operation of the x-ray source to prevent orreduce unwanted variations in the source output.

In front of one of the sensors 70 an amount of polymethyl methacrylate86 is placed to simulate an average thickness of soft tissue. In frontof other sensor 70 an amount of bone simulating material 84 is placed inan amount equivalent to that encountered in the spine or femur. Varioushydroxyapatite-epoxy mixtures are commercially available for bonesimulation in x-ray imaging. Therefore, a secondary detection systemwith a bone standard of known density and a soft tissue equivalentthickness is provided in this embodiment.

The signals from each sensor 70 can be used to compute the density ofthe bone internal standard as a function of time during the scan. Anydeviations from a constant density of this standard are due to changesin either the energy or intensity of the x-ray emission. Each value ofbone density computed in the patient scan corresponds to a computedvalue of the bone standard. Therefore, each computation of bone densityderived from a pair of high and low energy CCD frame acquisitions can becorrected or normalized by using the deviation from the density of theinternal standard. For example, if the value of the bone standard duringthe rectilinear scan deviated by plus 3% in a given area of the image,the computed bone density of the patient scan must be corrected by thatamount in this area. This internal reference approach can be used withall stationary and scanning embodiments described herein.

In conjunction with the above calibration approach, a number of strips72 (square rods) of bone simulating epoxy material, or aluminum ofequivalent x-ray absorption are placed under the table 73 which run inthe direction of the scan for the slit scan approach. Each linear striphas a different thickness or bone equivalent density. As the x-ray tubeand detector assembly is scanned over the area to be tested, each set ofrods are scanned and their density computed. The consistency of themeasured densities of these rods is used to ensure proper operation ofthe system. This set of standards can be placed anywhere from the x-rayexit port 80 to the edges 74 of the detector 76.

The imaging of radionuclide distributions in biological tissues orspecimens is a routine task performed in virtually all biomedicalresearch laboratories by the well established technique ofautoradiography. In this procedure, a thin slice of a specimen is placedin contact with photographic film thus allowing the radiation from thespecimen to expose the film. Subsequently, the film is processed bystandard chemical development techniques, manually, or by using anautomatic processor. Frequently, an intensifying screen is used in orderto enhance the absorption efficiency of the image receptor and for areduction in exposure time. Intensifying screens are especially usefulwhen images of relatively high-energy gamma or x-ray emissions arerecorded (20-200) keV. Also they can be useful for high energyelectrons.

Autoradiography produces images reflecting the biodistribution of aradionuclide and it has been established as a powerful tool in manybiomedical disciplines. Its major shortcomings relate to problems withquantization of the relative or absolute concentration of radionuclidein an area of interest. This difficulty arises from the non-linearity ofphotographic film typically used and in reciprocity law failure whenintensifying screens are used. Moreover, the development temperature,and in general, the condition of the processing chemicals have aninfluence on the film fog level and contrast. All these factors renderquantization a very difficult and time consuming task which becomesvulnerable to many uncertainties in quantitative autoradiography.Despite these problem, several investigators have digitized filmautoradiographs by using microdensitometers or video cameras for bothquantization and image enhancement.

In autoradiography, the image represents areas where the radiotracer hasbeen extracted. The anatomical information on the original tissue slideis not transferred with great detail in the autoradiograph. For properinterpretation, it is necessary to observe the tissue slide andautoradiograph side by side in order to correlate radiotracerdistribution with anatomy. Often it is necessary to superimpose theslide with the autoradiograph in order to identify the exact anatomiclocation of the radiotracer. In this process the accuracy in assigningan anatomic location to the tracer is severely compromised.

One of the most important problems with autoradiography is the longperiod of time required in order to expose the film. In mostapplications this time ranges from a few hours to several days, evenweeks in some cases. Therefore, the technician may have to wait for afew days in order to find out whether an exposure has to be repeated.

Autoradiography does not relate to in vivo imaging of radionuclidedistributions in humans or animals. Rather it relates to detectingradioactive distributions in excised samples. All available film-screenimage receptors have extremely low quantum efficiencies for most gammaemitters commonly used for this purpose. Moreover, the presence of alarge volume of tissue results in enormous amount of gamma ray scatterwhich will reach the image receptor and degrade the contrast and spatialresolution. The film-screen receptors do not have energy discriminationcapabilities, therefore scattered events cannot be rejected. The use ofa collimator to suppress scatter will result in a dramatic reduction ingeometric efficiency.

Thus the present invention, in its various embodiments, provides aneffective means for performing autoradiography by providing a compactdevice that performs the data acquisition for autoradiography quicklyand can superimpose both emission and transmission studies to correlatethe emission image with the anatomical features of the object underexamination. The embodiments described in connection with FIGS. 10 and11 below can be used to perform autoradiographic procedures.

Radionuclide imaging of humans and animals is performed on a routinebasis by using the Anger camera, most commonly referred to as a "GammaCamera". The gamma camera has a quantum efficiency in excess of 50% forthe most commonly used radionuclides and it has the capability ofdiscriminating scatter from primary photons by pulse-height analysis ofeach detected photon. The intrinsic spatial resolution of the gammacamera is approximately 3.5 mm. The total spatial resolution of thecamera, including the degradation due to its collimator, can vary from 5mm to 12 min. Modern gamma cameras can detect photons at the rate of25,000 counts per second (cps) without significant dead time losses. Athigher count rates, significant deviations are observed between true anddetected events. This is due to limitations inherent in the design ofboth the detector assembly and processing electronics.

The following presents a further embodiment relating to imaging ofradionuclide distributions in tissue samples and in vivo quantitativeimaging of humans and animals. This procedure employs a charge-coupleddevice to detect and process information to provide, in essence, acompact "gamma camera" using a highly sensitive stationary (or scanning)detector to conduct both emission and transmission studies at countrates up to 10⁶ of the object being examined.

Existing gamma cameras have limited spatial resolution, limitedcapability to perform in high count-rate conditions and it cannot beused to record x-ray transmission (radiographic) images with any degreeof acceptable detail to satisfy radiographic imaging standards.Therefore, the recording of a high quality radionuclide (physiologic)image and a radiographic (anatomic) image with the same detector foraccurate correlation of the physiologic and anatomic image remaindifficult. Where very high detail is necessary, the gamma camera isgenerally not capable of producing better than 5 mm resolution evenunder the most favorable conditions. Therefore, the imaging of smallparts of the body or imaging small animals like mice cannot be performedwith any reasonable detail using the gamma camera. This also applies forthe imaging of tissues containing radioactive materials.

The following procedures enable the acquisition of high detailradionuclide images and the option of combining them with the x-rayradiographic images with the same detector. This approach employs anovel acquisition scheme that enables imaging spectroscopy of gammarays, x-rays or nuclear particles by using a CCD. CCDs have beenemployed in the past without a scintillator for imaging spectroscopy ofvery soft x-rays, up to the energy levels of about 6-9 keV. However,above this energy, the CCD becomes virtually transparent to x-rays orgamma-rays. Generally, scintillators have not been used in conjunctionwith a CCD for imaging spectroscopy because it is believed that theconversion from gamma-rays to light will destroy the useful informationcarried by the interacting gamma-ray or x-ray. Therefore, imagingspectroscopy of gamma-rays or x-rays in nthe energy range of about 10keV to 2,000 keV with a CCD has not been explored. Also, alternating themode of operation from a counting, energy sensing detector to anintegrating detector for radionuclide and radiography, respectively,presents a useful procedure for imaging spectroscopy. Note, however,that the counting procedure can also be used in certain x-raytransmission measurements to measure the energy thereof.

When light interacts with the sensitive surface of the CCD, it generatesa charge which remains stored in the pixel where this interactionoccurred. As with previous embodiments the magnitude of the charge isdirectly proportional to the detected intensity of light. Each pixel isrepresented by its two-dimensional coordinates and by an intensityvalue. The energy required to produce an electron in the sensitivesilicon surface of the CCD is about 3.65 eV.

This value enables the determination of the energy of detected photonsif the system can either detect one photon at a time, or if the numberof the photons detected per pixel is known. This provides for imaging ofradionuclide distributions with a simultaneous measurement of the energyof the detected events. This procedure is termed "Imaging Spectroscopy"and provides a technique using gamma rays, beta-rays, and x-rays inconjunction with CCD technology.

The upper energy limit of soft x-ray imaging is between 5-10 keV. At 10keV, the quantum efficiency of a CCD is approximately 5% and itdiminishes rapidly at higher energies. The small fraction of the totalnumber of events interacting with the CCD will result in a high partialenergy transfer to the sensor with losses in proportion with the energyand the signal. Therefore, when the CCD is used as the primary detectorof high energy photons or particles, it is virtually unusable forperforming imaging spectroscopy. The following procedure provides highresolution imaging spectroscopy using a CCD that is suitable for manyapplications including position emission tomography and nuclear particleimaging.

A schematic of the device is shown in FIG. 10. An important component ofthis device is a CCD 98 with low readout noise, high charge transferefficiency and dark current levels. A CCD with less than 10electrons/pixel (RMS) readout noise is suitable for this purpose. Thedark current can be reduced to less than 0.6 electrons/sec at -40° C. bya compact thermoelectric cooler.

In one embodiment of this method, a thin scintillator 104 is used as theprimary detector of x-rays. One such scintillator can be a layer ofgadolinium oxysulfide or thallium activated cesium iodide or any of thecommonly available phosphors. The scintillator 104 is bonded to a fiberoptic faceplate 106 and the faceplate is bonded to an image intensifier96. The intensifier is bonded to a second faceplate 106 that is bondedto bundle 102. Optical bonding of this type is well established. Tofurther illustrate this embodiment the sensitive area of thescintillator 104, faceplates 106, image intensifier 96, fiber opticcoupler 102, and CCD 98 have, identical dimensions. Note that acollimator 94 can be mounted on the lead enclosure 100 and is usedduring the transmission study, and depending on its configuration, canalso be used during the emission study. Note that the collimator 94 canoptionally be removed during emission studies.

When an x-ray photon within the rays 14 interacts with the scintillator104, it produces light with intensity which is proportional to theenergy of the x-ray. This light is transported through the fiber opticfaceplate 106 and interacts with the CCD 98. The interaction of opticalphotons in each CCD pixel will produce a number of electrons in directproportion to the number of optical photons and to the energy of thedetected x-rays 14 or gamma-rays 92 that are produced by the isotopethat has collected in the lesion 90. Isotopes commonly utilized includeTC 99 m or I-125. The following example as a first order approximationof the expected energy resolution from the detector.

A 60 key x-ray interacts with the scintillator resulting in 3000 opticalphotons. Approximately one half of these photons are emitted in thedirection of the CCD. Assuming a Lambertian distribution of the emittedphotons from the screen, the transmission through the fiber optic plateis approximately 40%. Therefore, 600 optical photons will be arriving atthe CCD. The quantum efficiency of the CCD is approximately 40%,therefore only 240 photons will be detected in one pixel.

It can be shown that the energy resolution can be in the order of 10%which is approximately twice that attained with conventional NaI-crystalspectrometers at this gamma-ray energy.

FIG. 11 depicts an alternative embodiment in which a "pin hole"collimator 112 with shutter 110 is used in performing an emission studyof lesion 90 or any selected organ. The emission from the lesion ororgan impacts the scintillator 104, into housing 100, through the fiberoptic reducer 116, coupled to the intensifier 118, and than directed offmirror 124, lens system 120, and onto a cooled CCD 120.

This procedure produces radionuclide scintigraphy with spatialresolution in the order of about 1 millimeter or less, and transmissionimages with resolution in the order Of 0.2 millimeters. The spatialresolution and sensitivity of the detector will be selectable for bothemission and transmission modes via pixel binning. The detectoroperation will be selectable for pulse-height analysis or integration.For x-ray transmission imaging, the integrating mode of operation ispreferred. Note that during x-ray transmission imaging, the pin holecollimator will be removed. Emission imaging of thick tissues requires acollimator, either a multihole type or a pinhole collimator. Very thinspecimens can be imaged without a collimator by placing them very closeto the scintillator.

This camera has the capability of detecting very high count rates. Inconventional gamma cameras, each x-ray photon interaction occupies theentire scintillator and electronics for a period of time of 1 to 8microseconds after it is detected. In the present method, due to themultiple detectors, higher count rates can be handled due to themultiple detectors, and higher count rates can be handled without usinga scintillator with short decay time. Count rates up to 10⁶ counts persecond can be acquired with very low probability (less than 1%) ofdetecting 2 gamma ray events in one pixel when operating in thepulse-height analysis mode.

Note the scintillator can be bonded directly on the fiber optic bundlewithout the use of an image intensifier. Also, the scintillator can bebonded directly on the CCD without the use of a fiber optic bundle. Aframe transfer CCD is a preferred approach, but a full frame CCD can beused.

The following "shutter" methods can be used (a) a frame transfer CCD;(b) a gated image diode, or microchannel intensifier; or (c) a liquidcrystal shutter with very thin window or fiber optic window. The liquidcrystal shutter can be positioned between the fiber optic bundle and thescintillator.

Note that the system has applications for small animal imaging, skeletalimaging, monitoring of fracture healing, thyroid scintigraphy,Bremsstrahlung imaging of beta emitters within the body (radiationsynovectomy), intraoperative imaging probe, radionuclide angiography,small parts imaging, and pediatric nuclear imaging.

FIG. 12 illustrates in schematic form several methods that can be usedin performing quantitative imaging in accordance with the variousembodiments of the invention.

Note that one can use either a stationary source and detector to projectradiation 130, or a scanning source and detector assembly to scan theobject being examined 132.

Both stationary and scanning embodiments utilize a CCD detector thattransfers the detected information to a memory 140. The information canbe binned or processed 142 to accomplish various tasks. This processingcan include the application of software modules to correct fornon-uniformities in the source or collection components, or to identifyevents where light from one gamma-ray interaction has spread to a numberof neighboring pixels. Clusters of pixels with high intensity can beidentified as primary events and low intensity clusters can beidentified as scattered radiation and be eliminated by a filter.

Quantified information such as an intensity histogram (i.e., a pulseheight spectrum) can be generated 146 and a display of the object can begenerated 144 with the unwanted pixels removed.

After each set of data is produced in both the stationary and scanningembodiments, the conditions for operation can be modified 138 to producean image at a different energy level, to perform an emission ortransmission study, or to rotate the source and detector assemblyrelative to the object under study to produce three dimensional imagesor two dimensional images at different angles.

The emission and transmission studies can be displayed alone orsuperimposed. Due to the binning capability of the system a one to onecorrespondence exists between both emission and transmission images thatwas previously not possible. This high resolution image can be colorcoded to distinguish between the emission and transmission images.

Another preferred embodiment is illustrated in FIG. 13 where a fullframe or frame transfer cooled CCD 150 with a transparent scintillator152 bonded on the sensitive surface of the CCD, or to an imageintensifier 154, as shown. The scintillator 152 is preferably emittinganywhere from the UV blue to the red regions of the spectrum uponstimulation with x-rays or gamma-rays. The preferred scintillator is oneemitting in the green such as CsI (TI) or Cadmium tungstate, oralternatively a gadolinium based ceramic scintillator available fromHitachi Corporation. This scintillator has about twice the density ofsodium iodide or CsI(TI) and has higher efficiency. A fiber optic plate(straight or reducing) can be incorporated between the CCD andscintillator. Alternatively, an electrostatic image intensifier 154, orimage diode intensifier, can be incorporated between the scintillatorand the fiber optic plate. The scintillator 152 can be opticallytransparent plate or comprise a fiber optic array with fibers ranging indiameter from 0.006 mm to one or more millimeters. The thickness of theplate can be in the order of 0.5 mm to 5 mm.

Another preferred embodiment employs a CCD of the type described abovebut in conjunction with an electrostatic demagnifying image intensifier.The optical coupling of the CCD is accomplished by a fast lens at theoutput end of the image intensifier or by a fiber optic plate betweenthe output screen and the CCD.

The process of obtaining a desired image includes the initiation ofacquisition with the CCD for about one second or at a desired binningconfiguration, typically coarser than 2×2 pixels. Shorter acquisitiontime will be required for high count-rates and longer acquisition timeis tolerated for low count rates. The optimal acquisition time for aparticular application can be determined empirically by acquiring a fewtest frames and search for coincident events within individual pixels.Very short acquisition times (less than 1 millisecond) are easilyattainable by using a fast mechanical shutter, an electro opticalshutter, or by gating the image intensifier tube. This enablesacquisition with spectroscopy capability even at very high count-rates.Each acquisition "frame" will record from a few hundred to a fewthousand counts. After acquisition, each frame is stored in the computermemory for subsequent processing. Depending on the application, thetotal number of frames for a complete acquisition can vary, for example,from ten to a few hundred.

Each gamma-ray event in a given frame stored in the computer isrepresented by its x and y coordinates and by an intensity value (z)which is the number of electrons generated in this area of the CCD. Thez value is directly proportional to the energy of the gamma ray (orx-ray). The number of electrons generated from each interaction shouldbe confined to one pixel or group of binned pixels forming a"superpixel". In a significant percentage of interactions, the electronsgenerated from a single gamma-ray interaction can be split between twoor three pixels or superpixels. These split events form clusters in theimage matrix which can be easily identified by the computer software andassigned an x and y coordinate.

In one embodiment, as shown in the process flow sequences of FIG. 14,pulse height analysis uses the value of these neighboring pixels whichare summed to produce the z value for this gamma-ray event. Low z valuesrepresent gamma-rays which have been scattered and have lost a portionof their energy. These events are generally not desirable for inclusionin an image because they carry false position information. Therefore,the degree of rejection of each event can be decided by software on thebasis of the z value and a spectrum of the number of gamma-rays versusthe z value (energy) can be recorded. This filtering process can berepeated for each frame and all the frames can be added together to formthe final image. The operator can optionally go back to each originalframe, use a different z value threshold and reconstruct the final imageusing different filter parameters. Variations in the sensitivity of eachpixel or superpixel can be mapped and included in the counter for pixelby pixel corrections. The ability to discriminate different radiationsources measured simultaneously or sequentially includes defining filterparameters as selected energy threshold values or ranges.

In this radionuclide imaging technique, the degree of scatter rejectioncan be varied after the image acquisition in order to decide on theoptimal scatter rejection. This is not possible with the conventionalradionuclide imaging technology employing a gamma camera or arectilinear scanner. A gamma camera or rectilinear scanner is generallyincapable of detecting and processing high intensity x-rays which areemployed for high quality x-ray radiography.

If an image intensifier is not used, the scintillator can be in directcontact with the CCD. Alternatively, a fiber optic reducer can be usedbetween the CCD and the scintillator. Typical reduction ratios vary from1:1 to 6:1 although the present embodiment is not limited to theseratios. Therefore, for a 20 mm×20 mm CCD, and a 6:1 fiber optic reducer,the area of coverage will be about 120 mm. With a gated imageintensifier or a shutter, the CCD does not receive any signal during thereadout process. In a direct contact configuration, the use of frametransfer CCD as shown in FIG. 10 is preferred.

In applications utilizing x-ray transmission measurements a single frameis acquired for the recording of the x-rays emerging from the irradiatedbody of tissue. The CCD is operating in the integrating mode and eachpixel or superpixel which accumulates a charge which is proportional tothe total number of x-rays in this region without any energydiscrimination. The resulting radiographic image can be combinedelectronically with the radionuclide image to form an accuraterepresentation of both physiologic and anatomic information.

In the case of thin specimens examined in vitro a light source withwavelength ranging from the ultraviolet to near infrared can be used forthe transmission image in the integrating mode. In this approach, thelight shield in front of the scintillator is removed and the detector isplaced in an enclosure to shield it from ambient light.

The present invention can thus combine radionuclide emission imaging andx-ray transmission imaging (radiography) using the same area detectorwith spectroscopic capability in the gamma-ray imaging mode. This cameracan be operated utilizing both, the counting pulse-height analysis forgamma-ray imaging, and in the integrating or counting modes for x-raysubstantially transmission imaging. This enables exact superposition ofthe two images for accurate anatomic and physiologic imaging. Also, theoperator can change the energy threshold even after the radionuclideimage has been acquired. Thus, higher intrinsic spatial and energyresolution are provided than found in the conventional approaches.

FIG. 15 is a schematic illustration of one preferred embodiment of adual-energy bone densitometry system 200 in accordance with theinvention. An x-ray tube 12 emits x-rays 14 which pass through the x-raytransparent patient table 254 and into the patient (not shown). Thex-rays 15 which pass through the patient are directed through an x-raytransparent mirror 202 and strike a first scintillator screen 204. Thescintillator 204 reacts to low-energy x-rays and generates a lightpattern corresponding to the low energy x-ray pattern. The lightgenerated by the scintillator 204 propagates back to the mirror 202which reflects the light to the lenses 206. The lenses 206 couple theimage from the scintillator 204 to an image intensifier 208 havingmicrochannel plates 210. Alternatively, the image intensifier 208 can bea proximity-type intensifier without the microchannel plates 210. Thelight from the image intensifier 208 is received and detected by thedetector 212, which can be a CCD array, a CID array or an amorphoussilicon sensor. The detector 212 senses the image which corresponds tothe low-energy x-rays and generates an electronic representation of theimage in the form of pixel data.

High-energy x-rays pass through the scintillator 204 to an optionalx-ray filter 214. The filter 214 is preferably a copper filter whichblocks any remaining low-energy x-rays which pass through thescintillator 204. An optional light block filter 216 can also beincluded between the scintillator 204 and the x-ray filter 214 to blockany stray optical radiation emanating from the scintillator 204 fromreaching a second detector 220.

The high-energy x-rays from the filter 214 strike a second scintillator218 which is reactive to the high-energy x-rays to generate an opticalimage which corresponds to the pattern of high-energy x-rays. Theoptical image is received by a second detector 220, which can also be aCCD or CID array or an amorphous silicon image sensor. The seconddetector 220 senses the optical image and generates an electronicrepresentation of the high-energy x-ray pattern. An optional x-rayabsorbing fiber optic plate 222 can also be included between thescintillator 218 and the detector 220 to absorb any remaining x-rays andthus prevent them from interfering with the detector 220.

The system 200 of FIG. 15 can be used in either a scanning mode or astationary mode. In the scanning mode, the x-ray tube source 12 as wellas the detection system are moved continuously or in a stepping motionalong the region being examined. While the system scans the region, aseries of images are obtained having short exposure acquisition times.In the stationary mode, a single exposure is made of the entire regionbeing examined. Time delay integration (TDI) is used in which the CCDstores the total charge for each pixel during a selected x-ray exposureinterval. At the end of the x-ray exposure, the discrete representationin each pixel is readout by a CCD controller. Once the data is thusobtained, the comparative processing techniques of dual photonabsorptiometry can be used to determine quantitative densitymeasurements of the calcified material such as bone within the bodyregions exposed by the x-rays.

In the system 200 of FIG. 15, the image intensifier 208 can be omitted.In that configuration, to ensure that image data for the low-energyx-rays can be accurately collected, the detector 212 can be cooled toincrease signal-to-noise ratio.

FIG. 16 is a schematic diagram of another embodiment of a dual-energybone densitometry measuring system 300 in accordance with the invention.An x-ray tube 12 outputs x-rays 14 through x-ray transparent patienttable 254 and into the patient. X-rays 15 directed through the patientstrike a first scintillator 302 which is reactive to low-energy x-raysto generate an optical image of the low-energy x-ray pattern out of thepatient. The optical image is carried by a coherent fiber optic conduit304 to a CCD detector 306 which detects the optical image and generatesthe electronic representation of the low-energy x-ray pattern. The fiberoptic conduit 304 is preferably made of plastic optical fibers tofacilitate collection of the low-energy image. However, if the distancelabeled "x" is selected to be small enough, glass fibers can be usedinstead. The space labeled 310 is filled with a film material being thesame material as that of which the fibers are made.

The high-energy x-rays pass through the scintillator 302, the fiberoptic conduit 304 and the film material 310 and strike a second x-rayphosphor scintillator 312. The second scintillator 312 is reactive tohigh-energy x-rays and therefore generates an optical image whichcorresponds to the high-energy x-ray pattern. The optical imagegenerated by the scintillator 312 is detected by a second CCD array 314which generates the electronic representation of the high-energy x-raypattern. An optional copper or aluminum filter 316 can be inserted infront of the second scintillator 312 to absorb any remaining low-energyx-rays. Also, an x-ray absorbing fiber optic plate 308 can be insertedbetween the scintillator 312 and the CCD 314 to prevent x-rays fromimpinging on the CCD 314.

FIG. 17 is a schematic diagram of another embodiment of a dual-energybone densitometer measuring apparatus 400 in accordance with theinvention. The system 400 of FIG. 17 is the same as the system 300 ofFIG. 16 except that the coherent fiber optic conduit 304 in FIG. 16 isreplaced with a different conduit 404 in the system 400 of FIG. 17. Inthe conduit 404 of FIG. 17, the fibers are bent at approximate rightangles with small radii of curvature. As in the embodiment of FIG. 16,the fibers are either plastic or glass. Because of the different fiberbending in which the collected radiation is redirected from a firstoptical path onto a second optical path, the need for the film material310 shown in FIG. 16 is eliminated.

FIG. 18 is a schematic diagram of another embodiment of a bonedensitometer measuring apparatus 500 in accordance with the invention.In this embodiment, scintillator plates 505 and 507 are used to convertthe x-ray energy into optical energy. Once again, the x-ray tube 12directs x-rays 14 through the patient table 254 and the patient. Thex-rays 15 emanating from the patient first strike an anti-scatter grid502 which prevents scattered x-rays from reaching the detectors. Thex-rays then strike a first amorphous silicon image sensor 504 whichdetects low-energy x-rays and generates the data which indicates thelow-energy x-ray pattern. The low energy sensor 504 can be thinner thanthe high-energy sensor 508 to reduce the filtering requirements of thesystem. Also scintillator 505 can be thinner than scintillator 507 toimprove collection efficiency of the system. High-energy x-rays passthrough the first sensor 504 and then through a copper, tungsten,gadolinium or aluminum x-ray filter 506 which filters out low-energyx-rays. The high-energy x-rays then strike the second amorphous siliconimage sensor 508 which generates the data for the high-energy x-raypattern. The low-energy x-ray pattern data and the high-energy x-raypattern data are read out of the amorphous silicon image sensors 504 and508, respectively, by a detector controller 510.

FIG. 19 is a schematic diagram of an alternative detection structure 550which can be used with the dual-energy bone densitometry measuringapparatus 500 of FIG. 18. The lower layer of the structure 550 is ananti-scatter grid 552 used to prevent scattered x-rays from reaching thedetection structure 550. The next layer is a low-energy x-rayscintillator layer 554 which generates an optical image of thelow-energy x-ray pattern. An amorphous silicon image sensor 556 detectsthe optical image from the scintillator 554 to generate the data for thelow-energy x-ray pattern. A substrate layer 558 is formed over theamorphous silicon image sensor layer 556. The substrate layer 558includes a thinned central region 560. The thinned substrate 558provides for increased transmission to the second scintillator layer562. The second scintillator 562 is reactive to high-energy x-rays togenerate an optical image of high-energy x-ray pattern. The opticalimage is detected by a second amorphous silicon image sensor 564. Thestructure 550 is covered by a protective substrate 566, preferably madeof glass. A thin layer of lead can be formed on top of the glass toprevent propagation of x-rays beyond the structure 550. Preferredscintillators include C_(S) (+1), C_(d) WO₄, or gadolinium oxysulfide.

The amorphous silicon array sensors and the associated control andprocessing systems can utilize the binning and other processingcapabilities described elsewhere in the present application.Additionally, a plurality of such sensors can be combined to form asingle or dual array. The array can be linear, rectangular or squaredepending upon the particular application. The systems can be used inconjunction with a C-arm assembly where the C-arm 580 rigidly aligns thesource 586 and detector assembly 582 as shown in FIG. 20. The C-arm 580can also be used to rotate the source and detector about the patient ontable 584 as indicated at 588 to provide multidirectional viewing of theentire human skeletal structure including the hip and femur. Thuslateral spine imaging and quantitative analysis can be conducted usingthe present system. The detector assembly 582 includes a CCD sensor asdescribed herein in conjunction with a straight or angled fiber opticcoupler and scintillator. The detector assembly 582 can be scanned orstepped along axis 590 and axis 592 that is parallel to the spine of thepatient in order to provide a sequence of images for both quantitativeand qualitative analysis.

The detector assembly 582 can include various configurations describedelsewhere herein, including the examples illustrated in FIG. 21A, 21B,and 21C. In FIG. 21A a straight fiber optic coupler 602 opticallycouples the scintillator 604 to the CCD (or CID or amorphous) sensorarray 600. An optional cooler(s) 606 can be used in these examples. InFIG. 21B a fiber optic reducer 608 couples the scintillator 610 to thesensor array 600. A proximity type x-ray image intensifier andscintillator can replace scintillators 604 and 610. In FIG. 21C a dualsensor system includes sensors 600 and 612, fiber optic coupler 602,scintillators 618, 620, mirror 616, and lens 614. This system functionsin a manner similar to that described in connection with FIG. 15.

FIGS. 22A and 22B illustrate a preferred method of imaging in which theentire imaging field is composed of a series of slightly overlappingindividual images 620 that are acquired by a continuous scan or steppedimaging sequence along the rectilinear path 622. Dual energy tissue orbone density measurements can be accomplished by collecting data at twoenergies at each subfield 620. The x-ray source can be switched orfiltered as described previously to generate discrete energy peaks.

FIG. 23 illustrates a fan-beam system in which the x-ray source 586generates a fan shaped beam 640 that is detected by a detector system700. System 700 can include a scintillator, fiber optic plate or reducerfor each of a plurality of sensors 630 which are aligned in a lineararray to collect fan beam 640. Detector system 700 can use a lead slitcollimator 702 and can use CCDs, CIDs or a number of amorphous siliconsensors in configurations illustrated, for example, in FIGS. 21A-21C.

FIG. 24 illustrates another preferred embodiment 650 in which a patient654 is positioned on table 652. X-ray tube 656 directs fan-beam 660through a scanning slit collimator 658, the patient 654 and a secondscanning slit collimator 664. The radiation 660 then passes throughmirror 62 striking the scintillator 676. The scintillator emits lightthat is reflected by mirror 662 towards the sensor 672 as illustrated at674. Optional lead glass element 666 can be placed at any positionbetween the mirror 662 and the sensor 672. A lens 668 and cooler 670 canalso be employed, if necessary. Lead foil 678 can be used to line theenclosure 710 to reduce interactions between the scattered x-rays andthe sensor 672. The system can alternatively use a proximity type imageintensifiers in the x-ray path before the mirror.

FIG. 25 is a schematic diagram of a detection system 800 which can beused with the systems described above for dual-energy bone densitometrymeasurements as well as tissue and lesion imaging. The system 800 caninclude an enclosure 802 having an aperture 804 through which radiationsuch as x-ray beams 806 enter the system 800. In one embodiment, thex-ray beams 806 pass through an x-ray transparent mirror 808 and strikea first scintillating plate 809. The first scintillator 809 is reactiveto low-energy x-rays and produces an optical image corresponding to thelow-energy x-ray pattern. The optical image is projected back onto themirror 808 which reflects the image to lens 812. The lens 812 focusesthe light onto a first surface of a CCD array detector 814. The detector814 can include a proximity-type image intensifier to enhance imagedetection capabilities. An annular cooler (not shown) can also be placedaround the CCD detector 814 to cool the CCD and therefore improvesignal-to-noise ratio.

High-energy x-rays pass through the top scintillator 809 and strike thelower second scintillator 810 which is reactive to the high-energyx-rays to produce an optical image which corresponds to the high-energyx-ray pattern. The optical image is reflected by a second mirror 816 toa third mirror 818 which directs the light through a second focusinglens 820. The lens 820 focuses the light onto the back surface of theCCD detector 814. The back surface can also include a proximity-typeimage intensifier. In addition, the annular cooler, if present, coolsthe front and back surfaces of the CCD detector 814.

Thus, the system 800 of FIG. 25 produces increased sensitivity bysensing on opposite sides of a single thinned CCD detector 814. Spatialcorrelation between the two images which are fused to form a singleimage is greatly improved over the previously described embodiments withseparate detection surfaces since the relative locations of the imagedetection surfaces can be more precisely controlled. High and lowenergies can be detected on both sides.

The detection system 800 of FIG. 25 can include plural two-sided CCDdetectors to provide the system with a wide field-of-view along withcombined electronics and cooling. FIG. 25 illustrates a second detector834 and associated lenses 830 and 832. It will be understood that moredetectors and lenses can be added as needed.

The dual-energy configuration of the system 800 described abovefacilitates bone densitometry measurements as previously described.However, the system 800 can also be used for detecting and imaginglesions in patient tissue as described above. In that embodiment, thex-ray beams 806 are replaced with other types of radiation such as inthe visible or infrared ranges. The scintillators 809 and 810 and themirrors 808, 816 and 818 can be used to form images of the tissue to beformed at two different wavelengths on opposite surfaces of thedetectors 814 and 834.

A particular application of the methods and systems described herein fordetecting and imaging of soft tissue lesions includes digitalmammography using CCDs or similar type silicon-based detectors such asamorphous silicon type detectors described above. These systems are usedto detect lesions in the tissue including calcified material within thesoft tissue, that can indicate the need for more careful diagnosticprocedures and/or treatment of the patient. Slot-scanning approachesusing time-delay integration, where the CCD records continuously duringa scan as described herein, can be used for digital mammography.However, the continuous recording approach results in certain problems,particularly with artifacts due to the shear distortion of thefiber-optic plates which can be used with such an embodiment. Whileslot-scan approaches using the continuous record mode can be used, thequality of the images is less than ideal due to the distortion effects.

Other methods for scanning the breast include the dividing of the imagearea into four quadrants or even a greater number of segments. Everytime it is necessary to take multiple exposures of the breast, theassociated problems including increased exposure level and collectiontime can limit the variety of applications for which the system can beused. It is desirable, therefore, if one needs to acquire the image in astep-wise fashion, that there be no more than two, or at most, threeacquisition steps. If one uses a greater number of acquisition steps,the breast has to remain compressed for too long, thus causing extremediscomfort to the patient. Moreover, the x-ray tube power requirementsincrease significantly. The preferred method for digital mammographyapplications involving sequential multiple imaging is thus limited to atwo image acquisition process. This procedure involves directing x-raysfrom the source through the tissue to the stationary detector system forabout 0.2-5.0 seconds and preferably in the range of 0.5-1 second, thedetector system can then be moved to a second position while the firstimage is read out, then a second exposure is obtained and read out. Asmany as 2-5 million pixels can be read out in a time interval that isless than the exposure interval.

A problem associated with the two-dimensional array approach, is itscomplexity and cost. Although the tiling of 4×3 CCDs, for example, toform an array can be used for digital mammography, it is likely to betoo expensive for many common applications. This results from the costof the CCDs themselves, and with the problems associated with making aseamless joint to three or four sides of the CCDs.

Referring to FIGS. 26-32, the detector module 900 can consist of from 3to 5 CCDs in a first linear array 902 and another set of CCDs can bepositioned approximately 6 cm apart in a second linear array 904. Thisembodiment utilizes four CCD elements in each array. Each element caninclude a scintillator 906 and a tapered fiber optic plate 908. In theembodiment where the CCDs are replaced by amorphous silicon sensors, asingle strip silicon sensor can be substituted for each linear array inthese embodiments.

The first set of CCDs can be placed as closely as possible to the chestwall of the patient. The x-ray beam is collimated by using a double slotto provide two fan beams, each fan beam being directed onto a lineararray, thus only two areas are irradiated which correspond exactly toeach CCD group. After one x-ray exposure and acquisition, the x-raycollimator is translated in synchrony with both CCD banks which are alsotranslated to the next position. Another exposure is taken and thesignal is read out. A small amount of overlapping of the fields, about1-3 mm can be desirable. With the use of a micro-stepping translationstage the successive fields can be aligned to within a few microns withor without overlap. The images can then be joined and will besubstantially seamless with less than a 5-10 micron difference betweenthe region of the body and the joined images of the region.

The sensing surface does not have to be on a plane. As shown in FIG. 27,the CCDs 910 can be arranged on a curving or non-planar surface. This isan extremely important embodiment because it provides for the use ofstraight (non-tapering) fiber-optic plates 912 which dramaticallyreduces the cost and contributes to better image quality. Please notethat the CCDs can be cooled or non-cooled and can be operated in thepixel binned or non-binned mode. Additionally, an anti-scatter grid canbe used between the breast and the detectors. Each element 910 in thearray is generally equidistant from the x-ray source in order to reducedistortion across the entire field of view of the array. This arcedlinear array can be used for many different applications as describedelsewhere herein.

This approach is preferable as current manufacturers can readily makeCCDs which are buttable on two sides. It remains difficult and expensiveto make CCDs buttable on three or four sides. In the illustratedembodiment there are only six joints required between the CCDs, unlike alarge area cassette which has many more. A typical CCD for thisapplication can have an area of 6×6 cm but for economy reasons, one canuse a larger number of CCDs, such as 3×3 cm elements. For example, ifone uses a 3×3 cm device, each CCD linear array 902, 904 incorporateseight CCDs for a total of 16 CCDs. This can also be used to provide alarger area of coverage comparable to a standard large film cassette.FIG. 28 illustrates a four line separated array 916 covering a 24×30 cmplanar area with 10 of the 3×3 cm devices in each line 918. In FIG. 29,a partial cross sectional view of one of the lines 918 in FIG. 28,illustrates a preferred embodiment in which each CCD 914 is buttedagainst one or two adjoining CCDs in each line with each line coupled toa scintillator 915 and a fiber optic plate 917. The two-step acquisitionis preferable relative to the narrow slot-scanning approaches whichtypically use a slot width of about 1.5 cm, and the larger area imagingapproaches which are effective but can be extremely costly.

As illustrated in FIG. 30, an x-ray source 922 and a double or multipleslot collimator 924 can be used to generate and align the x-rays 928with the translating CCD modules. An actuator or motorized system 920 isused to translate both CCD arrays 902, 904 without altering the distancebetween the rigidly aligned CCD arrays. The system 920 can be connectedto a controller or personal computer as described previously so that theuser can control array position along the direction of translation 926,which in this embodiment, is towards or away from the chest wall.

As shown in FIGS. 31A and 31B, the arrays 902, 904 are positioned toimage two parallel regions 930, 932. The detectors 902, 904 are thentranslated from the first position to a second position to image andanalyze two further parallel regions 934, 936 to provide a full image ofcompressed breast 925. The relative spacing between the two lineararrays can also be controlled to increase or decrease overlap. Apreferred embodiment, however, retains the two spaced arrays in a rigidposition relative to each other. This particular embodiment moves thedetectors towards or away from the chest wall of the patient.

Shown in FIG. 32 is an embodiment 940 in which the array 916 of FIG. 28in which the direction of scan 942 is along the chest wall. Thecollimator 944 is also moved along the same axis as the array to directx-rays 928 onto the CCDs 948 and not onto the spaces 946.

While the invention has been particularly shown and described withreference to a preferred embodiment thereof, it will be understood bythose skilled in the art that various changes in form and details can bemade therein without departing from the spirit and scope of theinvention as defined by the appended claims.

The invention claimed is:
 1. An apparatus for examining soft tissue of apatient comprising:an x-ray radiation source emitting radiation which isdirected through the soft tissue; a scintillator receiving radiationtransmitted through the soft tissue and generating an optical signalcorrelated with the received radiation and directing the optical signalalong a first optical path; an optical surface receiving the opticalsignal along the first optical path and redirecting the optical signalalong a second optical path; a binning charge couple device (CCD) thatis optically coupled to the optical surface to receive the optical signalong the second optically path at a plurality of pixels and generate anelectronic representation of the soft tissue; and a CCD controller thatis electrically connected to the CCD such that the CCD controller binscharge from separate pixels for readout to a data storage device.
 2. Theapparatus of claim 1 further comprising a display for displaying animage of the soft tissue.
 3. The apparatus of claim 1 furthercomprising:a display for displaying an image of the soft tissue; and adata processor receiving the electronic representation of the softtissue and manipulating the image of the soft tissue on the display. 4.The apparatus of claim 1 further comprising a cooler in thermalcommunication with the CCD to cool the CCD.
 5. The apparatus of claim 1further comprising an image intensifier coupling the optical image tothe CCD.
 6. The apparatus of claim 1 further comprising processing meansfor determining the density of the soft tissue wherein the soft tissuecomprises a lesion.
 7. The apparatus of claim 1 further comprising aframe holding the radiation source and the CCD in fixed relation to eachother.
 8. The apparatus of claim 1 wherein the CCD is uncooled.
 9. Theapparatus of claim 1 further comprising a processor that combines datastored by groups of adjacent pixels of the CCD to generate theelectronic representation of the soft tissue.
 10. The apparatus of claim1 wherein the radiation source is located above a support surface onwhich the patient can be positioned and wherein the optical surface ispositioned between the support surface and the scintillator.
 11. Theapparatus of claim 1 wherein the CCD performs time delay integration togenerate the electronic representation of the soft tissue.
 12. Theapparatus of claim 1 wherein the CCD comprises an array of pixels largerthan 512×512 pixels.
 13. The apparatus of claim 1 further comprising alens between the optical surface and the CCD focusing the optical signalonto the CCD.
 14. The apparatus of claim 1 wherein the first and secondoptical paths are at a right angle to each other.
 15. The apparatus ofclaim 1 wherein the CCD further comprises a sensor that is binned duringcharge readout such that rectangular groups of pixels are combined. 16.The apparatus of claim 1 wherein the binnable CCD comprises 2×2 pixelbinning CCD.
 17. An apparatus for examining soft tissue of patientcomprising:an x-ray radiation source emitting radiation which isdirected through the soft tissue; a scintillator receiving x-rayradiation transmitted through the tissue and generating an opticalsignal correlated with the received radiation and directing the opticalsignal along a first optical path; a mirror transparent to the radiationtransmitted through the tissue located between the scintillator and theradiation source, said mirror receiving the optical signal from thescintillator and reflecting the optical signal along a second opticalpath; a binning charge coupled device (CCD) that is optically coupled tothe mirror to receive the optical signal along second optical path at aplurality of pixels and generate an electronic representation of thesoft tissue; a CCD controller that is electrically connected to the CCDsuch the CCD controller bins charge from separate pixel readout to adata storage device; and a lens between the scintillator and the CCDalong the second optical path, the lens focusing the optical signal ontothe CCD.
 18. The apparatus of claim 17 further comprising a display fordisplaying an image of the soft tissue.
 19. The apparatus of claim 17further comprising:a display for displaying an image of the soft tissue;and a data processor receiving the electronic representation of the softtissue and manipulating the image of the soft tissue on the display. 20.The apparatus of claim 17 further comprising a cooler in thermalcommunication with the CCD to cool the CCD.
 21. The apparatus of claim17 further comprising an image intensifier coupling the optical image tothe CCD.
 22. The apparatus of claim 17 further comprising processingmeans for determining the density of the soft tissue wherein the softtissue comprises a lesion.
 23. The apparatus of claim 17 furthercomprising a frame holding the radiation source and the CCD in fixedrelation to each other.
 24. The apparatus of claim 17 wherein the CCD isuncooled.
 25. The apparatus of claim 17 further comprising a processorthat combines data stored by groups of adjacent pixels of the CCD togenerate the electronic representation of the soft tissue.
 26. Theapparatus of claim 17 further comprising CCD means for performing timedelay integration to generate the electronic representation of the softtissue.
 27. The apparatus of claim 17 wherein the CCD comprises an arrayof pixels larger than 512×512 pixels.
 28. The apparatus of claim 17wherein the first and second optical paths are at a right angle to eachother.
 29. A method of examining soft tissue in a patientcomprising:positioning the soft tissue of the patient between an x-rayradiation source and a scintillator; directing radiation emitted by theradiation source through the soft tissue; receiving radiation with thescintillator that is transmitted through the soft tissue and generatingan optical signal correlated with the received received radiation anddirecting the optical signal along a first optical path; redirecting theoptical signal along second optical path receiving the optical signalalong second optical path a plurality of pixels of a binning chargecoupled device (CCD); binning charge from separate of the CCD forreadout with a CCD controller that is electrically connected to the CCD;generating a first electronic representation of the soft tissue; androtating the CCD relative to the soft tissue and generation a secondelectronic representation of the soft tissue.
 30. The method of claim 29further comprising displaying an image of the soft tissue.
 31. Themethod of claim 29 further comprising:displaying an image of the softtissue; and receiving the electronic representation of the soft tissueand manipulating the image of the soft tissue on the display.
 32. Themethod of claim 29 further comprising cooling the CCD.
 33. The method ofclaim 29 further comprising coupling the optical image to the CCD withan image intensifier.
 34. The method of claim 29 wherein the soft tissuecomprises a lesion.
 35. The method of claim 29 further comprising thestep of holding the radiation source and the CCD in fixed relation toeach other.
 36. The method of claim 29 wherein the CCD is uncooled. 37.The method of claim 29 further comprising the step of combining datastored by groups of adjacent pixels of the CCD to generate theelectronic representation of the soft tissue.
 38. The method of claim 29further comprising the step of locating the radiation source above thepatient whose tissue is being examined.
 39. The method of claim 29further comprising the step of performing time delay integration togenerate the electronic representation of the soft tissue.
 40. Themethod of claim 29 wherein the CCD comprises an array of pixels largerthan 512×512 pixels.
 41. The method of claim 29 further comprisingfocusing the optical signal onto the CCD with a lens between the opticalsurface and the CCD.
 42. The method of claim 29 wherein the first andsecond optical paths are at a right angle to each other.
 43. The methodof claim 29 further comprising performing serial and parallel binning ofpixels of the CCD.
 44. The method of claim 29 further comprising binningpixels during charge readout of the CCD.
 45. A method of examining softtissue of a patient comprising:directing radiation emitted by theradiation source through the soft tissue; directing radiationtransmitted through the soft tissue through a mirror transparent to theradiation; receiving the radiation with a scintillator that istransmitted through the mirror and generating an optical signalcorrelated with the received radiation and directing the optical signalback toward the mirror along a first optical path; redirecting theoptical signal along a second optical path with the mirror; focusing theoptical signal along the second optical path; receiving the focusedoptical signal along the second optical path at a plurality of pixels ofa charge coupled device (CCD); binning charge from separate pixels ofthe CCD, the CCD being electrically connected to a CCD controller; andperforming readout of the CCD to form a binned electronic representationof the soft tissue.
 46. The method of claim 45 further comprisingdisplaying an image of the soft tissue.
 47. The method of claim 45further comprising:displaying an image of the soft tissue; and receivingthe electronic representation of the soft tissue and manipulating theimage of the soft tissue on the display.
 48. The method of claim 45further comprising cooling the CCD.
 49. The method of claim 45 furthercomprising coupling the optical image to the CCD with an imageintensifier.
 50. The method of claim 45 wherein the soft tissuecomprises a lesion.
 51. The method of claim 45 further comprising thestep of holding the radiation source and the CCD in fixed relation toeach other.
 52. The method of claim 45 wherein the CCD is uncooled. 53.The method of claim 45 further comprising the step of combining datastored by adjacent pixels of the CCD to generate the electronicrepresentation of the soft tissue.
 54. The method of claim 45 furthercomprising the step of locating the radiation source above the patientwhose tissue is being examined.
 55. The method of claim 45 furthercomprising the step of performing time delay integration to generate theelectronic representation of soft tissue.
 56. The method of claim 45wherein the CCD comprises an array of pixels larger than 512×512 pixels.57. The method of claim 45 wherein the first and second optical pathsare at a right angle to each other.
 58. A method of imaging soft tissuein a patient comprising:providing an x-ray radiation source such thatradiation emitted by the source is transmitted through a patient's softtissue onto a scintillator; providing an integrated silicon circuitsensor having a two dimensional array of pixel elements that detectlight from the scintillator that is emitted in response to radiationfrom the x-ray source; positioning a region of soft tissue the patienton a support surface; directing x-ray radiation through through theregion of the patient's soft tissue onto the scintillator which emits aspatial intensity pattern of light that is detected by the sensor, thespatial intensity pattern being coupled to the sensor with anon-intensified optical system; binning charge from separate pixelelements of the sensor for readout with an electronic controller; andforming an image of the soft tissue from the binned representation. 59.The method of claim 58 wherein the step of providing a pixellated sensorcomprises providing a binnable sensor.
 60. The method of claim 58further comprising providing a data processor connected to the sensor,the data processor having a memory for storing the discrete electronicrepresentation.
 61. The method of claim 58 further comprising performingpulse height analysis on the discrete electronic representation.
 62. Themethod of claim 58 further comprising providing a fiber optic couplerbetween the scintillator and the sensor.
 63. The method of claim 58further comprising forming the image in less than 60 seconds afterdirecting the x-ray radiation through the patient.
 64. The method ofclaim 58 further comprising simultaneously irradiating the entire regionof the patient with x-ray tube that is stationary relative to thepatient.
 65. The method of claim 58 further comprising forming an imagehaving a resolution of at least about 1 mm.
 66. The method of claim 58further comprising providing a sensor having a two dimensional array ofMOS diodes.
 67. The method of claim 58 further comprising performingserial and parallel binning of pixels of the sensor.
 68. The method ofclaim 58 further comprising binning pixels during charge readout of thesensor.